The present invention generally relates to apparatus and methods for magnetic resonance imaging (MRI), also known as nuclear magnetic resonance imaging (NMRI). More particularly the present invention relates to methods and apparatus for decreasing magnetic resonance data acquisition times including time for reconstructing the image wherein the data is acquired and the image is reconstructed in parallel. The present invention also relates to other methods and MRI systems and apparatus related thereto.
Magnetic resonance imaging (MRI) is a technique that is capable of providing three-dimensional imaging of an object. A conventional MRI system typically includes a main or primary magnet(s) that provides the background magnetic field Bo, gradient coils and radio frequency (RF) coils, which are used for spatial encoding, exciting and detecting the nuclei for imaging. Typically, the main or primary magnet(s) are designed to provide a homogeneous magnetic field in an internal region within the main magnet, for example, in the air space of a large central bore of a solenoid or in the air gap between the magnetic pole plates of a C-type magnet. The patient or object to be imaged is positioned in the homogeneous field region located in such air space. The gradient field and the RF coils are typically located external to the patient or object to be imaged and inside the geometry of the main or primary magnet(s) surrounding the air space. There is shown in U.S. Pat. Nos. 4,968,937 and 5,990,681, the teachings of which are incorporated herein by reference, some exemplary MRI systems.
In MRI, the uniform magnetic field B. generated by the main or primary magnet(s) is applied to an imaged object by convention along the z-axis of a Cartesian coordinate system, the origin of which is within the imaged object. The uniform magnetic field Bo being applied has the effect of aligning the nuclear spins, a quantum mechanical property of macroscopic particles comprising the imaged object, along the z-axis. In response to RF pulses of the proper frequency, that are orientated within the XY plane, the nuclei resonate at their Larmor frequencies. In a typical imaging sequence, the RF signal centered about the desired Lamor frequency is applied to the imaged object at the same time a magnetic field gradient Gz is being applied along the z-axis. This gradient field Gz causes only the nuclei in a slice with a limited width through the object along the XY plane, to have the resonant frequency and to be excited into resonance.
After excitation of the nuclei in the slice, magnetic field gradients are applied along the x-axis and y-axis respectively. The gradient Gx along the x-axis causes the nuclei to precess at different frequencies depending on their position along the x-axis, that is, Gx spatially encodes the precessing nuclei by frequency (i.e., frequency encoding). The y-axis gradient Gy is incremented through a series of values and encodes the Y position into the rate of change of the phase of the precessing nuclei as a function of gradient amplitude, a process typically referred to as phase encoding.
The quality of the image produced by the MRI techniques is dependent, in part, upon the strength of the magnetic resonance (MR) signal received from the precessing nuclei. For this reason an independent RF coil is placed in close proximity to the region of interest of the imaged object in order to improve the strength of the received signal. Such RF coils are sometimes referred to as local coils or surface coils.
There is described in U.S. Pat. No. 4,825,162 a surface coil(s) for use in MRI/NMRI imaging and methods related thereto. In the preferred embodiment of that invention, each surface coil is connected to the input of an associated one of a like plurality of low-input-impedance preamplifiers, which minimizes the interaction between any surface coil and any other surface coils not immediately adjacent thereto. These surface coils can have square, circular and the like geometries. This yields an array of a plurality of closely spaced surface coils, each positioned so as to have substantially no interaction with all adjacent surface coils. A different MR response signal is received at each different one of the surface coils from an associated portion of the sample enclosed within the imaging volume defined by the array. Each different MR response signal is used to construct a different one of a like plurality of different images then being combined, on a point-by-point, basis to produce a single composite MR image of a total sample portion from which MR response signal distribution was received by any of the array of surface coils.
In the case of MRI phased-array coils, coils are de-coupled by two mechanisms; any adjacent pair of coils are de-coupled by overlapping and non-adjacent coils are de-coupled by combination of matching circuits and low impedance pre-amplifiers. The use of a phased array RF coils or surface coils with a tuned and matched circuit including low impedance pre-amplifiers have been used for de-coupling as well as a mechanism for improving the signal-to-noise ratio (SNR) and field of view (FOV). In this regard, it should be understood that the term xe2x80x9ccouplingxe2x80x9d refers to the coupling of an MR signal in one coil to an adjacent coil such that the signal being outputted by the adjacent coil is a combination of the MR signal detected by the adjacent coil and the coupled MR signal. Consequently, the image from the adjacent coil would be distorted to some degree. Although the tuned and matched circuit including low impedance pre-amplifiers has been effective from the standpoint of improving SNR and FOV, such circuitry becomes ineffective when both the number of coils and the coil density is increased. In other words, as the spacing between adjacent coils and between adjacent portions of a coil is decreased signal coupling is increased irrespective of the tuned and matched circuits.
Although there are a variety of spatial encoding methodologies or techniques being implemented, the most popular method used in commercial MRI scanners is two dimensional Fourier transform (2DFT) encoding in which a two-dimensional spatial plane (e.g., XY plane) is encoded with both frequency and phase of the MR signals. Typically during one data acquisition, only a one dimensional time-domain signal is obtained and thus 2DFT encoding requires repeating the data acquisitions sequentially to achieve a pseudo second dimension of the time domain signals. The second dimension of the spatial information is encoded into the phase component by repeating the data acquisition with different phase encoding gradient strengths (i.e., varying Gy to create the other pseudo-time dimension. In this technique, each slice of the imaged object is in effect divided into a multiplicity of layers in the y-direction or axis corresponding to the number of pixels in that direction (e.g., 128, 256). The number of pixels in turn is representative of the desired image resolution, in other words the higher the resolution the higher the number of pixels. In addition, the x- direction scanning process or the data acquisition is repeated by sequentially and repeatedly stepping through each of these y-direction layers. Because the resolution of the time-domain signals depends on the number of repetitions of the data acquisitions, and the repetition rate is limited by the MR relaxation times; a higher resolution image takes a longer time.
MR imaging has proven to be a valuable clinical diagnostic tool in a wide range of organ systems and pathophysiologic processes. Both anatomical and functional information can be gleamed from the MR data, and new applications continue to develop with each improvement in basic imaging technique and technology. For example, the ability to image and evaluate increasingly finer anatomical details have resulted with technological advances yielding improved achievable spatial resolution. Also, the technological advances allowing for fast imaging sequences has resulted in reduced imaging times such that many moving structures can be visualized without significant motion artifacts.
Often, however, there is tradeoff between spatial resolution and imaging time because higher resolution images require a longer imaging time. This balance between spatial and temporal resolution is particularly important in cardiac MR, where fine details of coronary artery anatomy must be discerned on the surface of a rapidly beating heart. Thus, a high-resolution image acquired over a large fraction of the cardiac cycle, will be blurred and distorted by the motion of the beating heart.
One technique for decreasing imaging time has concentrated on increasing speed of sequential scanning of K-space and thus acquisition of MR data by reducing the intervals between scanned lines. Because it has appeared difficult to significantly better efficiency of such conventional fast imaging, other fast imaging schemes have been proposed which schemes use simultaneous data acquisition in multiple RF coils. Such other schemes are described in detail in U.S. Pat. No. 5,910,728, the teachings of which are thus incorporated herein for such purpose.
Two recent methods, the Simultaneous Acquisition of Spatial Harmonics (SMASH) imaging in the time domain or k-space and the Sensitivity Encoded (SENSE) imaging in the frequency domain, changes such sequential data acquisition into a partially parallel process by using a phased array, thereby reducing the scan time as compared to the sequential data acquisition technique. In these two techniques, it is recognized that the data sampled below the Nyquist sampling rate can be recovered if the sensitivity profiles of the phased array detectors can provide enough spatial information to either interpolate the data in the time domain or unwrap the data in the frequency domain.
The time domain method or the SMASH method recognizes the equivalence between phase-encoding with MRI gradients and the composite spatial sensitivity inherent in the detectors. The SMASH method uses a numerical fitting routine to, among other things, interpolate a decimated number of phase-encoding steps and thus, achieve reductions in scan time. Although this numerical approach was instrumental in demonstrating the original SMASH concept, such a methodology ignores or does not recognize the underlying analytical relationship between the weighting factors for the composite harmonics, the image field-of-view (FOV), the spacing of the detectors, the harmonic orders, and the sensitivity profiles of the detector coils.
The SMASH method contains the following steps. First, sensitivity profiles of each of the phased array coil elements are derived from a separate data acquisition by MRI. Second, by using numerical fitting and computation, such as minimum least square or gradient-descent algorithms, the coefficients of linear combinations that compose the optimal sensitivity harmonics from the phased-array coils are numerically derived. Third, using composite harmonics to interpolate decimated phase encoding steps, the sampling rate is restored to the Nyquist frequency. Fourth, a Fast Fourier Transform (FFT) of the composite harmonics gives the non-aliasing MR image. There is described in U.S. Pat. No. 5,910,728 a conventional implementation of the SMASH methodology that utilizes a numerical gradient-descent fitting routine to generate a set of spatial harmonics from the sensitivity profile of a multi-channel array of MRI detectors to achieve multi-fold reductions in the gradient phase-encoding steps.
Also, the recursive numerical fitting routine for harmonic generation must simultaneously accommodate phase errors introduced by the individual detectors in the array. The phase errors imparted by each detector arise from the difference in the phase of the transverse magnetization generated at a point in space, as it is detected by each of the coils in the array due to the detector coils different locations in space. These phase errors may cause serious problems for generating spatial harmonics if they are not dealt with properly. Phase errors introduced by fixed or time-dependent acquisition delays, flow or motion, etc, on the other hand, will be the same for each detector coil so that their effect on the generation of harmonics will be insignificant although they can be the cause of image artifacts. In the implementation of conventional phased-array MRI using the SMASH technique, the numerical fitting technique used to generate the harmonics also includes compensating for spatial phase errors.
As to the SENSE method, and like the SMASH method, the maximum aliasing fold that can be unwrapped is limited to the number of elements or detectors making up the phased array coils. The SENSE method also requires precise sensitivity maps of all the detector coils in the phased-array. Thus, the SENSE method needs a large amount of preparation before reconstruction. Also, another problem with this method is that in locations where the sensitivities of multiple coils cannot be distinguished from each other, the SENSE reconstruction will fail.
A problem with such parallel data acquisition and reconstruction techniques is that conventional MRI phased array coils are unable to deploy a large number of coils due to the limitations imposed by both their loop structure and the de-coupling requirements for the mutual induction between the elements. Because the number of coils in the conventional phased array corresponds to the maximum decimation factor for reducing the number of phase encoding steps, existing phased array designs significantly limit the potential for parallel spatial encoding using the SMASH technique.
It thus would be desirable to provide an improved technique, method or procedure for spatially encoded MRI using the sensitivity profiles of an array of detectors, including a method for correcting the phase errors of the signals arising from the different detectors. It also would be desirable to provide MRI apparatuses or systems embodying or utilizing such a technique or procedure so as to allow the efficient combination, processing and reconstruction of the acquired decimated parallel MRI data. In addition, it would be desirable to provide techniques, methods or procedures that could be extended for use with MRI systems or apparatuses that deploy large numbers of coils or detector elements. Further, it would be desirable to provide techniques, methods or procedures that are adaptable so as to simulate spatial harmonic generation and evaluate conditions that introduce error and distortion of composite signals as a mechanism, for example, to maximize the number of useable harmonics for image reconstruction. As compared to the prior art techniques, apparatuses, and systems utilizing a numerical fitting procedure, the improved technique or the apparatuses or systems embodying such a technique would be analytical.
The instant invention is most clearly understood with reference to the following definitions or the terms used in equations unless otherwise separately defined.
w(x, y) represents the spin density distribution weighted by the relaxation times T1 and T2.
f(x, y) is the sensitivity profile of each individual detector.
xcfx86(x, y) represents the phase error introduced by each detector.
C(ky, n) forms a set of weighting coefficients
xcex3 is the gyromagnetic ratio
The present invention features an analytical methodology for spatially encoded MRI using the sensitivity profiles of an array of detectors, including a methodology for correcting the phase errors of the signals from the different detectors comprising the array. As compared to prior art techniques such an analytical methodology advantageously yields an analytical transform providing inter alia a quantitative relationship between the weighting coefficients of the composite signals, the detector geometry, the sensitivity profile, the image field-of-view (FOV). Also, such an analytical methodology advantageously yields harmonic order and the detector index. Additionally, and as compared to prior art techniques, such an analytical methodology advantageously provides a mechanism for removing or dealing with space-related phase errors of the detectors and analytically restoring the phase coherence among the signals from the array of detectors. Further, such a method also advantageously avoids the potential burden and cost of using hardware to correlate the phases of the signals in multiple receivers or the prior art reconstruction technique of the root-of-the-sum-of-the-squares. Also featured are other methods or techniques related thereto as well as apparatuses or systems embodying any of a number of the herein described methods, procedures or processes according to the present invention.
According to one aspect of the present invention, there is more particularly featured a method for correcting spatially-related phase errors and restoring phase coherency among MR signals received substantially independently from a plurality of RF detectors forming a plurality of receive channels, the detectors being configured so as to form an array of detectors (i.e., a phased array of detectors). The method includes applying a Fourier transform to the independently received MR signals so as to convert each received MR signal of each receive channel from k-space to image domain and applying a Hilbert transform to the magnitude of each converted signal in the image domain to generate the signal""s minimum phase. Such a method also includes applying an inverse Fourier transform to convert each signal with magnitude and minimum phase from the image domain back to k-space. In this way, the resultantly converted k-space signals are substantially free of spaced related errors.
According to another aspect of the present invention there is featured a method for parallel spatial encoding an MR image data that is frequency-encoded by MR gradient and sensitivity-encoded by the plurality of RF detectors of an RF detector array. The method includes applying an analytical transform function to generate weighting coefficients for a given spatial harmonic order and a given detector index. The numerator of the analytical transform is a complex exponential function having an exponent that is the product of 2xcfx80, the spatial harmonic order, the RF detector index, RF detector spacing and the reciprocal of the field-of view. The denominator of the analytical transform is a Fourier transform of the sensitivity profile of the detector. The method also includes generating linear combinations of the MR image data that is frequency-encoded and sensitivity-encoded to generate a set of spatial harmonics that can encode spatial frequencies. The method further includes applying at least a one-dimensional Fourier transform to a k-space data set in which spatial frequency dimensions are fully encoded, thereby resulting in an MR image of an observed slice of an object being observed.
The method further comprises demodulating the modulation of high order harmonics being generated by said applying an analytical transform function. The step of demodulating includes applying another Fourier transform to composite spatial harmonics to determine frequency and amplitude of a modulation component and adding another component, having same frequency and amplitude but opposite phase as the modulation component, to the generated high order spatial harmonics. In this way, the modulations of the high order harmonics are substantially eliminated.
According to yet another aspect of the present invention there is featured a method for imaging with an apparatus having multi-coil arrays and various MRI pulse sequences whose phase-encoding gradient increment is increased b-fold, resulting in a b-fold reduction in the total number of phase-encoding steps. In this method, the decimated raw data from each channel of the detector array are saved for reconstruction. This method includes acquiring a reference image or a sub-set of image data to obtain the sensitivity profiles of the array coils and acquiring the partial gradient phase-encoded signals from the phased array coils for each slice of the object being observed. The phases of the signals from the phased array coils is synchronized. More specifically, the signals are synchronized using a Fourier-Transform-Hilbert Transform (FT-HT) phase correction process of the present invention. The method also includes generating the harmonics to replace the phase-encoding steps and combining the harmonics by interleaving them to form a set of fully encoded raw data. The method further includes applying a 2-dimensional (2D) Fourier Transform to the fully encoded raw data to reconstruct the image.
Other aspects and embodiments of the invention are discussed below.